Medical imaging is one of the most useful diagnostic tools available in modern medicine. Medical imaging allows medical personnel to non-intrusively look into a living body in order to detect and assess many types of injuries, diseases, conditions, etc. Medical imaging allows doctors and technicians to more easily and correctly make a diagnosis, decide on a treatment, prescribe medication, perform surgery or other treatments, etc.
There are medical imaging processes of many types and for many different purposes, situations, or uses. They commonly share the ability to create an image of a bodily region of a patient, and can do so non-invasively. Examples of some common medical imaging types are nuclear medical (NM) imaging such as positron emission tomography (PET) and single photon emission computed tomography (SPECT), electron-beam X-ray computed tomography (CT), magnetic resonance imaging (MRI), and ultrasound (US). Using these or other imaging types and associated machines, an image or series of images may be captured. Other devices may then be used to process the image in some fashion. Finally, a doctor or technician may read the image in order to provide a diagnosis.
A conventional “block” detector utilizes a 2×2 array of photomultiplier tubes (PMTs) to identify a gamma-ray scintillation event within a pixilated scintillation crystal block by computing the position of the incident gamma-ray from a logical combination of PMT output signals. FIG. 1 shows a basic block detector design using four PMTs 110 (A, B, C and D) to compute the location of a scintillation event within crystal block 100.
In conventional timing readout for PET-based PMT systems, wherein coincidence must be detected between a pair of oppositely traveling gamma-rays produced from the annihilation of a positron, the total energy signal from the PMT array is used for signal timing purposes. This is shown in FIG. 2, wherein the anode (A) outputs of the four PMTs A, B, C and D are summed to obtain an energy signal E (also sometimes denoted as Z). The summing amplifier (not shown) typically has to be a very high bandwidth, low noise amplifier with low input capacitance.
FIG. 2 illustrates the HV resistor network of the prior art configuration, wherein, using conventional “Anger” logic (see U.S. Pat. No. 3,011,057 to Anger, issued Nov. 28, 1961 and incorporated herein by reference) the total energy of a gamma-ray event is calculated as E=A+B+C+D, and the X and Y spatial coordinates are calculated as:
  X  =                                          (                          A              +              B                        )                    -                      (                          C              +              D                        )                          E            ⁢                          ⁢      and      ⁢                          ⁢      Y        =                            (                      A            +            C                    )                -                  (                      B            +            D                    )                    E      
Generally, the signal timing is obtained from the energy signal E through known constant fraction discriminator (CFD) circuits. The timing resolution will be determined by the rise-time, signal-to-noise ratio, and the input capacitance of the signal E. It is difficult to obtain a good timing signal from summing the anode outputs A, B, C and D because the anode outputs must be split into two branches, one for obtaining the energy signal E and the other for obtaining the position coordinates X and Y.
Baseline shift creates another problem found in AC-coupled circuitry correlated with signal count-rate. Baseline shift may be caused by charge buildup in isolating capacitors. AC-coupling systems usually overcome this problem by implementing baseline-restore software or hardware, which may increase costs.
Therefore, there exists a need in the art to improve detector timing performance, reduce baseline shift and simplify timing readout electronics.